1. Field of the Invention
The present invention relates to improvements in ultrasonic imaging; and, more particularly, to a high speed adaptive ultrasonic phased array imaging system. The system employs an image sharpening process which maximizes the average brightness of image texture within a selected region of interest by varying the phased array scan data of array elements for the image lines within the region.
2. The Prior Art
Ultrasonic imaging has been extensively applied in virtually every medical specialty in the form of pulse echo B-mode tomography (See Wells, 1977). This modality displays echoes returning to the transducer as brightness levels proportional to echo amplitude. The transducer is mechanically or electronically translated or steered in one dimension. The brightness levels are displayed versus echo range and transducer position or orientation, resulting in cross sectional images of the object in the plane perpendicular to the transducer face.
Ultrasound B-scan systems may incorporate piston-like piezoelectric transducers fabricated to produce a fixed focus achieving diffraction limited spatial resolution in the lateral direction. These are mechanically scanned to obtain the tomographic image. Alternatively, segmented array transducers are now used in many medical ultrasound imaging devices (See Wells, 1977). A number of types of transducers may be used. In one type of transducer, an annular phased array of transducer elements are arranged in a bulls-eye pattern (See Melton et al, 1978) and incorporated into a mechanically driven imaging system. Annular arrays enable focusing of the ultrasound beam during both transmit and receive operations, including receive mode dynamic focusing over a long depth of focus, by properly timing or "phasing" the transmit pulses and selectively delaying the receive mode echoes.
Another type of transducer, a sequential linear phased array (see Wells, 1977), operates by sequentially activating groups of piezoelectric elements for the transmit-receive process in the on-axis direction. Each group of elements produces a single image line resulting in a rectangular image format. As with the annular array, transmit focusing and receive mode dynamic focusing are achieved by proper timing of the transmit signals and delaying receive mode echoes. A third type of transducer, a sectored linear phased array consists of a single group of transducer elements which is not only focused by also steered over a sector angle in transmit and receive (Tx and Rx) by properly timing the transmit signals and receive mode echoes (see von Ramm et al, 1983).
In each type of phased array imaging system, the correct timing relations for transmit and/or receive modes are pre-calculated and incorporated into the imaging device as hard wired circuitry or data contained in the software memory of a digitally controlled system. In phased array imaging systems in the prior art, this timing or phasing data is determined by assuming propagation of ultrasound pulses through a homogeneous tissue medium with a uniform velocity of sound, usually 1540 m/sec. The assumption of a constant velocity of sound in the body is also the design basis in all ultrasound scanning systems for converting round trip pulse-echo time of flight into target range in the image.
Unfortunately, this simplest model of all human tissues is not valid. The body is actually composed of inhomogeneous layers of differing tissues (fat, muscle and bone) with bumps and ridges of varying thicknesses and different acoustic velocities. These layers intervene between the transducer and the internal organ of interest. The propagation velocity of ultrasound varies from approximately 1470 m/sec in fat to greater than 1600 m/sec in muscle and nervous tissue to as much as 3700 m/sec in bone (see Goss et al). If an incorrect average velocity is chosen, B-scan imaging is known to result in an image range error and compound scan registration errors for all ultrasound systems. A method to minimize these errors by offering a selectable average velocity has been previously described by Jellins and Kossoff (1973) for a water coupled ultrasound breast scanner.
Under the assumption of a uniform tissue medium of constant velocity, the presence of inhomogeneous tissues can also result in image artifacts, range shifts, geometric distortions, broadening of the transducer beam pattern which degrades the ideal diffraction limited lateral resolution, and increased side lobes which reduce the signal to noise ratio in the image. These problems occur in all types of pulse echo ultrasound systems to some degree. A worst case situation occurs in adult cephalic imaging through the skull layer where the problems are so severe that ultrasound imaging through the intact adult skull has been almost totally abandoned. Undesirable refraction effects from fat/muscle interfaces have also been noted in several clinical studies of abdominal ultrasound (see Muller et al, 1984) as well as in studies which simulate conditions in ultrasound breast imaging (see Davros et al, 1985).
The adverse effects of inhomogeneous nonuniform tissue layers have been analyzed by several investigators primarily in terms of unknown phase aberrations associated with the inhomogeneities introduced across the transducer aperture. Attempts have been made to overcome these aberrations using various signal processing techniques.
The present inventors have modeled the inhomogeneous tissue layers of varying thickness as planar layers of an assumed uniform thickness (see U.S. patent application Ser. No. 016,427 filed Feb. 19, 1987, incorporated herein by reference; and Smith et al, 1986). The refraction effects of these layers are then corrected in the sectored phased array data resulting in some improvement to image quality. In the case of adult cephalic imaging, the inventors have also described an on-line multiplacative receive mode signal processing technique (see Smith et al, 1976; and Smith et al, 1975) which improves image quality for the limited case of an intervening tissue layer whose aberration function is symmetric about the transducer center.
In recent years, attempts have been made to measure the phase aberration profile in front of the transducer for the purpose of phase correction. The present inventors applied this technique to ultrasound imaging (see Smith et al, 1978) by measuring the aberration function in front of the transducer array using the arrival time at each array element of signals from a single independent point source or point reflector. The sector scan phased array delay data for steering and focusing in the entire image was then updated using this measurement of the aberration function.
A similar technique was proposed by Hirama et al (1982) using echoes from one or a few discrete resolved point scatterers in the focal plane of an ultrasound array. In this case, the aberration function is measured along a single image line by varying the relative phase delays between two array elements at a time so as to maximize a quality factor relative to signal strength for the strongest single target in the desired focal plane. The measurement of the aberration function was used to correct the phased array data of an ultrasound C-scan in which the image plane is parallel to the transducer face. Hirama and Sato (1984) later realized the inadequacy of assuming discrete resolved point scatterers in a complex object and proposed a new technique for structured targets in which the spatial frequency contribution of the target is removed from the aberration measurement by extensive measurements of object spatial frequencies followed by image reconstruction via inverse Fourier transform.
It should be noted that the phase aberration function across the transducer aperture can change as a function of position or steering angle. The aberration is constant only over a region of target angle and target range known as the iso-planatic patch (IPP) (see Fried, 1974). Thus correction of phase aberration over an entire image using a phase aberration measurement or an image sharpening technique based on a single target angle and focal range usually is not valid. The size of the iso-planatic patch is inversely related to the severity of the transducer phase aberration, transducer aperture size and frequency. It has been demonstrated that the IPP for conventional phased array adult cephalic imaging is approximately 10.degree. (see Miller-Jones, 1980). It is anticipated that the IPP for conventional abdominal scanning is larger.
Another technique to restore diffraction limited spatial resolution degraded by aberrations was proposed by Muller et al, 1974, for astronomical optical telescopes which all use spatially incoherent radiation. An image quality parameter such as the intensity to a power integrated over a region containing a star or bright spot is maximized by varying the positions of an array of mirrors comprising the telescope, thus restoring image resolution. Subsequent analysis of this image sharpening process for spatially coherent radiation such as radar, microwave imaging, and ultra-sound imaging was carried out by Steinberg et al (see Attia, 1984; Steinberg, 1986; and Steinberg et al, 1986). Their conclusion was that image sharpening using intensity quality factors was not valid for spatially coherent imaging except by using a single point target. They have proposed another technique to change a spatially coherent imaging device into a spatially incoherent system by means of transmitter location diversity (also known as "spatial compounding" in the medical ultrasound art). The spatial incoherence would then enable the use of the image sharpening technique of Muller and Buffington (see Muller et al, 1974) over complex objects to compensate for transducer phase aberrations.
The present inventors have discovered that these conclusions in the prior art are incorrect regarding compensation for phase aberration via image sharpening in spatially coherent medical ultrasound imaging devices. Since 1978, the understanding has developed (see Burckhardt, 1987; and Wagner et al, 1983) that the texture of medical ultrasound images of tissue consists primarily of a random speckle interference pattern resulting from the phasor summation of echoes from a large number of fine scatterers within the transducer resolution cell (see Burckhardt, 1987); and Wagner et al, 1983). The echoes from these particles exhibit phases uniformly distributed over 0 to 2 .pi. radians. Although the image brightness of an individual speckle is a random process, the means image brightness and variance over an area is predictable.
The inventors have recently demonstrated that individual speckle spots change unpredictably from bright points to null as the phase function or aberration changes across the transducer aperture. However, the inventors have also demonstrated that the average image brightness of speckle in a region of interest is predictably decreased by transducer phase aberrations. Thus, an individual speckle in the image cannot be used as an image sharpening target. However, the average brightness of many speckles over a region of interest can be used as a quality factor in an image sharpening process for a phased array ultrasound scanner. When the discrete specular targets are also present in a region of interest, image sharpening using mean brightness in that area is less effective. On the other hand, if the uniform image texture in the region contains significant contributions from ordered unresolved scatterers in tissue, the image sharpening technique still performs quite satisfactorily.
Applicants have recently become aware of yet another method in the art for phase aberration correction described by M. O'Donnell and S. Flax, although Applicants are unaware as of this writing of any publications or patents describing this method. In this method for a region of interest in an ultrasound speckle image, a cross correlation function is calculated between two channels N and N+1 of a phased array system. The phased array scan data between these two elements is varied until a maximum is achieved in the cross-correlation function. The process is then continued with element N+2 versus N+1. This method using a cross-correlation relies on a product, i.e., multiplication, rather than an integral or a sum, as in the instant invention.
The phase correction technique of the instant invention was achieved totally independently of the method of O'Donnell et al and is significantly different. The method of O'Donnell et al using a product operation (cross-correlation) is more sensitive to slight phase differences between channels N, N+1. As such it is more sensitive to noise than the instant invention in the environment of low signal to noise ratio of typical ultrasound imaging. The cross-correlation function is a much more complex operation than a simple integral and is therefore more time consuming, thus limiting its application to high speed systems. Furthermore, the method of O'Donnell et al must be performed on the radio-frequency echoes prior to envelope detection. These signals range from 3.5 to 20 MHz and thus require high speed analog-to-digital (A/D) conversion. The instant invention operates on the echo signal after envelope detection which reduces the frequency requirements to approximately 1 MHz and is much more naturally and easily adapted to conventional ultrasound scanners. Finally, the method of O'Donnell et al requires A/D conversion of every phased array channel whereas the instant invention requires A/D conversion of only the summed signal from all the phased array channels.
There is a final important concept. In 1983, it was demonstrated that the average size of the speckle interference pattern is predictable is measured by the statistical parameter, the normalized auto-convariance function or its Fourier Transform pair, the noise power spectrum (see Wagner et al, 1983). The average speckle size Sc in the lateral direction can be defined by the full width half maximum (FWHM) of the normalized auto-convariance function (ACVF). For a fully developed speckle with no transducer aberration, ##EQU1## where .lambda. is the transducer wavelength, Z is the transducer focal range, and D is the transducer aperture length. It has recently been discovered that the average speckle size is also affected by the transducer aberrations. The main lobe width of the speckle normalized auto-convariance function decreases in the presence of phase aberrations. Thus, the speckle size offers another independent parameter for adaptive processing to compensate for transducer phase aberrations.
In summary, the present inventors note the following differences between a number of the above-discussed techniques and the present invention.
Jellins and Kossoff (1973) describe a water coupled breast scanner which features a single selectable average velocity. They make no attempt to compensate for tissue layers of different velocities within the body to overcome phase aberration.
Smith et al (1987 & 1986) model inhomogeneous tissue layers of varying thickness and velocity as planar layers of varying velocity. They correct for refraction errors due to the planar layers (minor effect) but make no attempt to correct for other principal sources of phase aberration such as layers of varying thickness.
Smith et al (1976, 1975) described an on-line multiplicative receive model signal processing technique which would correct for an intervening tissue layer whose aberration function is symmetric about the transducer center. This is a rare occurrence and their scheme cannot correct for a generalized aberration.
Phillips et al (1975), Smith et al (1978) and MillerJones (1980) described a method to measure the aberration function of an intervening tissue layer using a point reflector in the body or a point source on the opposite side of the body from the transducer. Once the aberration function has been measured, the phased array scan data can be corrected to eliminate the effects of the aberrator. However such point targets are rare in the body so that this technique would seldom be useful. The use of a second external transducer which must be aligned with the phased array would be clumsy and complex and may not be practical.
Hirama et al (1982) described a similar technique using echoes from a single target or the strongest target of a group of discrete scatterers. But groups of discrete resolved targets seldom exist in the body so the technique is seldom practical.
Hirama et al (1984) later realized the inadequacy of assuming discrete resolved point scatterers in a complex object and proposed a new technique for structured targets in which the spatial frequency contribution of the target is removed from the aberration measurement by extensive measurements of the target spatial frequencies followed by image reconstruction via inverse Fourier transform.
A similar technique to restore diffraction limited spatial resolution degraded by aberration was proposed by Muller et al (1974) for astronomical optical telescopes which all use spatially incoherent radiation. An image quality parameter such as the intensity to any power is integrated over any sized region for discrete targets. This may include a single point target such as a single star or any arbitrary small region of an extended source such as planet or galaxy which has a bright spot. The critical difference between the instant invention and Muller et al (1974) is that for the spatially coherent radiation of medical ultrasound, an arbitrarily small region including only a few image points will not suffice. The texture of medical ultrasound images consist primarily of a random interference pattern called speckle which acts as multiplicative noise modulating the real objects in an ultrasound image. Thus individual image points or speckles in the image of an organ do not correspond to real objects within the organ. During a process of phase correction the brightness of individual speckles will oscillate, and the integral of intensity to any power over an arbitrary small region will also oscillate. The improvement of the instant invention over the prior art is that the inventors combined the knowledge of previous prior art schemes with the recent understanding of ultrasound speckle to arrive at the method which requires an integral over a statistically significant region of speckle to achieve an increase in integrated brightness as the phase correction proceeds to the solution.
A further difference between the instant invention and Muller et al is that for a uniform object such as a planet with no bright spot, the method of Muller et al will not converge. However, a uniform object in medical ultrasound produces the image speckle pattern described above. Thus the brightness integral over a sufficiently large region will serve as an adequate quality factor.
Attia (1984), Steinberg (1986) and Steinberg et al (1986) also realized the futility of applying the technique of Muller et al (1974) to spatially coherent radiation of medical ultrasound. Their solution, however, was to change ultrasound images into spatially incoherent images by the technique of transmitter location diversity (known as spatial compounding) in medical ultrasound. If ultrasound images approach spatial incoherence, the arbitrary size of the ROI of Muller et al (1974) may suffice.
The following is a list of articles cited above.
Attia E. H., "Phase synchronizing large antenna arrays using the spatial correlation properties of radar clutter," Ph.D. Dissertation, University of Pennsylvania, 1984. PA1 Burckhardt C. B., "Speckle in ultrasound B-mode scans," IEEE Trans. Son. Ultrason., 25(1) pp. 1-6, Jan. 1978. PA1 Davros W. J., Madsen E. L. and Zagzebski J. A., "Breast mass detection by US: A phantom study", Radiology 156, 773-775, 1985. PA1 Fried D. L., Isoplanatic aspects of predetection compensation imagery. Report #TR-131, Optical Science Consultants, Yorba Linda, Calif., 1974. PA1 Goss S. A., Johnston R. L. and Dunn F., "Comprehensive compilation of empirical ultrasonic properties of mammalian tissues," J. Acoust. Soc. Am. 64(2), 423-457. PA1 Hirama M. Ikeda O. and Sato T., "Adaptive ultrasonic array imaging system through an inhomogeneous layer," J. Acoust. Soc. of Amer. 71(1), 100-109, 1982. PA1 Hirama M. and Sato T., "Imaging through an inhomogeneous layer by least-mean-sqaure error fitting," J. Acoust. Soc. Amer. 75(4), 1142-1147, April, 1984. PA1 Jellins J. and Kossoff G., "Velocity compensation in water-coupled breast echography," Ultrasonic 11, 223-226, 1973. PA1 Melton H. E., Jr. and Thurstone F. L. "Annular array design and logarithmic processing for ultrasonic imaging," Ultrasonic in Medical and Biology 4, 1-12, 1987. PA1 Miller-Jones S. M., "Automated arrival time correction for ultrasonic cephalic imaging." Ph.D. Thesis, Duke University, Durham, N.C., 1980. PA1 Muller N., Copperberg P. L., Rowley V. A., Mayo J., Ho B. and Li D.K.B., Ultrasound Med. 3, 515-519, 1984. PA1 Muller R. A. and Buffington A., "Real time correction of atmospherically degraded telescope images through image sharpening," J. Opt. Soc. Amer. 64(9), 1200-1210, 1974. PA1 Phillips, D. J., Smith, S. W., von Ramm, O. T., and Thurstone, F. L., "Sampled aperture techniques applied to B-mode echoencephalography", Acoustical Holography, Vol. 6, N. Booth, ed., 103-120, Plenum Press (New York and London, 1975). PA1 Shattuk, D. P., Weinshenker M. D., Smith S. W. and von Ramm O. T., "Explososcan: a parallel processing technique for high speed ultrasound imaging with linear phased arrays," J. Acoust. Soc. Amer. 75(4), 1273-1282, 1984. PA1 Smith S. W., Miller E. B., von Ramm O. T. and Thurston F. L., "Signal processing techniques to improve B-mode echoencephalography," Ultrasound in Medicine, D. N. White, Editor, Plenum Press, 1975, Vol. 1, 405-414. PA1 Smith S. W., Phillips D. J., von Ramm O. T. and Thurstone F. L., "Real time B-mode echoencephalography," Ultrasound in Medicine 11, White & Barnes, Editors, Plenum Press, 1976, 373-382. PA1 Smith S. W., Phillips D. J., von Ramm O. T. and Thurstone F. L., "Some advances in acoustic imaging through skill," Ultrasonic tissue characterization II, M. Linzer, ed., NBs Pub. #525, June, 1978, 209-218. PA1 Smith S. W., Trahey G. E. and von Ramm O. T., "Phased array ultrasound imaging through planar tissue layers" , Ultrasound in Medicine and Biology. 12(3), 229-243, 1986. PA1 Steinberg B. D., "Distortion correction by image feedback control" Valley Forge Research Center Quarterly Report #49, 54-58, 1986. PA1 Steinberg B. D., and Subbaram H. M., "Self calibration of phased array using the Muller Buffington Theorem and transmitter location diversity" Valley Forge Research Center Quarterly Report #50, 31-45, 1986. PA1 von Ramm O. T. and Smith S. W., "Beam steering with linear arrays," IEEE Transactions on Biomedical Engineering, BME-30 438-452, August 1983. PA1 Wagner R. F., Smith S. W., Sandrik J. M., and Lopez H., "Statistics of Speckle in Ultrasound B-Scans," IEEE Trans. Son. Ultrason. 30(3) pp 156-163, May 1983. PA1 Wells P. N. T., Biomedical Ultrasonics, Academic Press: London and New York, 1977.
The following patents disclose ultrasonic phased array systems having means for variable phasing of the array elements.
U.S. Pat. No. 4,566,459 to Umemura et al discloses a device and method for measuring the acoustic velocity inside a body for ultrasound imaging. The ultrasound diagnosis system determines whether the reflection signal of an object at a predetermined position is in-focus or out-of-focus and knows the actual acoustic velocity of the object from the assumed velocity when the reflection signal is in-focus.
U.S. Pat. No. 4,279,157 to Schomberg et al discloses a phased array scanning system which pre-calculates the transit time for acoustic beams, and is said to subsequently correct the variations of acoustic refractive indexed by comparing the calculated time against measured values.
U.S. Pat. No. 4,140,022 to Maslak and U.S. Pat. No. 4,149,420 to Hutchinson et al generally relate to ultrasound imaging.